Apparatus and method for tissue regeneration

ABSTRACT

An apparatus for tissue regeneration is provided. The apparatus comprises means for generating at least one laser pulse comprising a wavelength; and means for directing the at least one laser pulse onto a tissue surface of a human or animal body, wherein the means for generating comprises control means to ensure that a sum of the pulse energies of the at least one laser pulse is selected so that the corresponding fluence on the tissue surface heats the tissue surface up to a maximal temperature T max  between 70° C. and a tissue boiling temperature T b . Further, the means for generating of the apparatus are adapted so that a delivery time t ed  of the at least one laser pulse (during which the second half of the pulse energy is delivered) is sufficiently short so that, given the wavelength and thus a corresponding penetration depth δ of the at least one laser pulse, a thermal exposure time t exp  of the tissue surface is shorter than 900 microseconds. Here, the thermal exposure time t exp  of the tissue surface is defined as a time interval in which the temperature of the tissue surface is above T o +(T max −T o )/2, wherein T o  defines the initial temperature of the tissue surface, before the laser pulse arrives.

This application claims priority to European Patent Application No.18172363.6 filed on May 15, 2018, which is hereby incorporated herein byreference.

FIELD OF THE INVENTION

The present invention relates to an apparatus and a method for tissueregeneration. The goal of the treatment is the rejuvenation of thetissue by stimulating its regenerative potential.

TECHNICAL BACKGROUND

Nowadays, lasers are used for treating a wide variety of disorders andcosmetic conditions, in particular the rejuvenation of skin, mucosa orother tissue. Various treatments have been classified as rejuvenationtreatments, such as, for example, skin resurfacing and skin tightening,as well as vascular, pigment and acne treatments. In what follows we usethe expression tissue rejuvenation more narrowly to describe a treatmentwhere the tissue (e.g., one or more layers of the tissue) is thermallyinjured in a non-ablative or minimally ablative and/or reversiblemanner. In the subsequent wound healing process, the tissue experiencesa certain degree of regeneration. Such tissue regeneration is performednot only for cosmetic reasons, i.e. to make the patient's skin appearyounger, but has also been found effective in treating various disorderssuch as incontinence, atrophy, prolapse, vaginal laxity or analconditions.

In what follows, the terms “tissue” or “human tissue” will be used torepresent both human and animal tissue. Similarly, the terms“rejuvenation” and “regeneration” will be used interchangeably.

The human body consists of several types of the tissue, among them theepithelial tissue and the connective tissue (cf. FIG. 1 ). Epithelialtissue covers most of internal and external surfaces of the body and itsorgans. Epithelial tissue forms boundaries between differentenvironments, and nearly all substances must pass through an epithelium.In its role as an interface tissue, epithelium accomplishes manyfunctions, including protection of the underlying tissues from physicaltrauma and the detection of sensation. The principal cell type that isfound in epithelia are keratinocytes that generate biomoleculesnecessary for the stability and resistance of the epithelial layer tomechanical stress. In skin, the very thin top layer of the epitheliumconsists of dry dead cells (cf. FIG. 1 ).

The type of epithelial tissue that lines various cavities in the body,such as the mouth or vagina, and covers the surface of internal organsis the mucous membrane or mucosa. Another type of the epithelial tissueis the epidermis which covers the skin surface.

The epithelial tissue and the underlying connective tissue are separatedby the basement membrane, a thin, fibrous, extracellular matrix oftissue. The connective tissue lies below the epithelial tissue and helpsto hold the body together. The cells of connective tissue includefibroblasts, adipocytes, macrophages, mast cells and leucocytes.Fibroblasts are the most common cells of connective tissue. A fibroblastis a type of cell that synthesizes the extracellular matrix and collagenand plays a critical role in wound healing.

In the case of the human skin, the connective tissues are part of thedermis which is a layer of skin between the epidermis and subcutaneoustissues. The oral and vaginal connective tissue is termed laminapropria. Similarly, the lumen of the urethra is surrounded by epitheliumwhich, in turn, is surrounded by collagen-rich connective tissue and amuscle layer.

Since it is the connective tissue which is responsible for holding theskin, vagina and other organs together, previous rejuvenation techniqueshave typically focused on regenerating the connective tissues such asthe dermis or the lamina propria. Such a regeneration mechanism is basedon injuring the connective tissue, in order to induce a reactiveinflammatory response which results in an increase of the biosyntheticcapacity of fibroblasts and other cells. This leads to thereconstruction of an optimal physiologic environment, the enhancement ofcell activity, hydration, and the synthesis of collagen, elastin and HA(hyaluronic acid). Typically, the inflammation is achieved by deliveringheat to the connective tissue resulting in an increased temperature (ΔT)of the target tissue.

As will be explained below, assuming a single biochemical process thethermal damage to the tissue has as an approximately exponentialdependence on the exposure temperature T and a linear dependence on theexposure time (t_(exp)) and can be approximately described by theArrhenius integral relation. This relation predicts that, for a oneorder of magnitude decrease of the exposure time, the treatmenttemperature can be increased by approximately five degrees. Therefore,shorter exposure times are safer and also allow more intense treatmentsbecause of the higher acceptable exposure temperature.

When using a laser to create a laser pulse, the duration and the shapeof the resulting thermal exposure pulse within the tissue typically doesnot follow the duration and the shape of the delivered laser pulse. Thisis, because the volume of the heated tissue is typically relativelylarge, and the major mechanism by which this large volume cools down isthe relatively slow diffusion of the deposited heat into the surroundingunheated tissues. It is therefore the thermal diffusion rather than thetemporal pulse width of the delivered laser pulse which typically setsthe lower limit for the achievable exposure time. Typical cooling timesof the connective tissues are in the order of seconds or longer. Thislimits the treatment temperatures for the regeneration of the connectivetissue to about 45 to 70° C.

It should be noted, however, that, in spite of the fact that the allowedregeneration temperatures are relatively low, it is a considerablechallenge to heat up the deeper lying connective tissue to thesetemperatures. This is, because the delivered laser pulse must firsttraverse the epithelial layer located above the basement membrane,before the laser pulse can reach the fibroblasts. This means that thedelivered pulse energy may be predominantly absorbed by the superficiallayers. Accordingly, the maximally allowed temperatures for theepithelial layer limit the temperatures which can be generated withinthe connective tissues. Thus, the physician is often faced with atrade-off between using enough energy for an effective therapy andstaying within the damage thresholds for the superficial tissue. To acertain degree, this limitation may be overcome by applying externalcooling of the superficial layer prior or during the treatment. However,in many applications, such as when treating narrow body cavities (forexample vagina, urethra or anus), such a cooling may not be clinicallydesirable or technically feasible.

It follows from the above that the rejuvenation devices and methods aretypically designed for thermally bypassing the superficial layer to thelargest extent possible, in order to be able to directly thermallyactivate the deeper lying fibroblasts in a safe manner. This approach isbased on the conventional wisdom that an injury of the connective tissuewhich leads to an inflammatory response is essential for promotingcollagen production. Further, it is believed that an inflammation of theconnective tissue is needed to attract cells to the site of injury suchas neutrophils and macrophages which, in turn, release growth factorsand cytokines responsible for repair.

However, it should be noted that not only the fibroblasts but also thesuperficially located keratinocytes are involved in the wound healingprocess. In particular, it is well-known that keratinocytes recruit,stimulate and coordinate the actions of multiple cell types involved inhealing. Besides, keratinocytes and fibroblasts communicate with eachother via double paracrine signalling loops (known as cross talk ordynamic reciprocity) which coordinate their actions to restore normaltissue homeostasis after the wounding. In response to paracrinesignalling from keratinocytes and inflammatory cells, fibroblastssynthesize collagen and promote cross-linking to form an extracellularmatrix.

Accordingly, an apparatus and a method are desired which focus on anintense and safe thermal activation of the superficially locatedepithelia instead of bypassing the superficial layer for reversiblyinjuring the deeper located connective tissue.

Further, there has been a prejudice in the prior art that regenerationmethods should not heat up the treated tissue to more than 70° C., sinceotherwise significant tissue damage would occur due to proteindenaturalization.

Terminology

Before we turn to the present invention, we first define some quantitieswhich are important for explaining the present invention. The basicsetting of the present invention is that a laser pulse having a finitewidth impinges on the surface of a tissue. On the tissue surface, thelaser beam gives rise to a spot S, wherein the spot S can be defined asthe region of the tissue surface within which 90% of the total energy ofthe laser pulse is delivered to the tissue (i.e., the spot S correspondsto the smallest region on the tissue surface so that, during thedelivery of the laser pulse, 90% of the total energy of the laser pulseis delivered to this region).

Referring to FIG. 2 , the energy of the laser pulse partially transmitsthrough the tissue surface and passes the tissue down to a penetrationdepth δ, wherein the penetration depth represents the inverse of theabsorption coefficient within the tissue at the wavelength of the laserbeam.

As the energy of the laser pulse is mainly absorbed within thesuperficial layer of the tissue which has the depth δ, this superficiallayer of the tissue is rapidly heated up from the initial tissuetemperature T_(o) to a maximum temperature T_(max).

Further, it is noted that the tissue temperature increase ΔT (t) duringa laser intensity pulse I(t) (in W) is not proportional to I(t) but tothe cumulative laser energy E(t) (in J), defined by:

$\begin{matrix}{{E(t)} = {\int_{0}^{t}{{I(t)}{dt}}}} & (1)\end{matrix}$

The pulse energy E_(o) of the laser pulse is defined by E_(o)=E(t=t_(p))where the intensity pulse duration (t_(p)) represents the time duringwhich 95% of the total energy of the laser pulse is delivered to thetissue. Depending on the application, other percentage values of thetotal energy may be used.

The energy delivery time t_(ed) of the laser pulse is then defined ast_(ed)=t_(p)−t_(eh), where the time t_(eh) represents the time when thecumulative energy E(t) reaches half of the pulse energy,E_(o)/2=E(t=t_(eh)). Thus, the energy delivery t_(ed) is the time spanduring which the second half of the pulse energy E_(o) is delivered.

The upper part of FIG. 3 illustrates the intensity curve I(t) of a laserpulse, wherein, as mentioned above, the pulse duration t_(p) is the timeduring which 95% of the total energy of the laser pulse is delivered tothe tissue.

The lower part of FIG. 3 illustrates the corresponding temperature curveof the irradiated tissue surface. Starting with a temperature T_(o) (thetemperature of the tissue before the laser pulse impinges on the tissuesurface), the laser pulse heats up the surface of the irradiated tissueto a maximal temperature T_(max).

It is noted that the heating phase of the tissue lasts until time t_(p),i.e., until the end of the energy delivery by the laser pulse. Then, thecooling phase begins during which the superficial layer of the tissuecools down in the absence of any external energy delivery to the tissue(see the lower part of FIG. 3 ). The cooling is predominantly caused bythe fast diffusion of the heat from the thin superficial tissue layer tothe underlying deeper tissue layers, and to a smaller extent also by anatural convection into the surrounding air or other surrounding media.It should be appreciated that diffusion-mediated cooling also takesplace during the heating phase.

Defining the various quantities shown in FIG. 3 more precisely, we startwith the thermal exposure time t_(exp) which is the time span for whichthe treated tissue is exposed to elevated temperatures. Referring to thelower part of FIG. 3 which shows the temporally varying tissue surfacetemperature, we define the thermal exposure time t_(exp) as the timedifference t₂−t₁ between those two points on the temperature curve atwhich the temperature difference ΔT(t)=T(t)−T_(o) reaches half itsmaximum value, i.e., where ΔT(t₁)=ΔT(t₂)=(T_(max)−T_(o))/2 holds. Thus,the thermal exposure time t_(exp) corresponds to the FWHM (full widthhalf maximum) value of the temperature curve. In other words, thethermal exposure time t_(exp) is the time interval in which thetemperature of the tissue surface is above T_(o)+(T_(max) T_(o))/2

Further, the energy exposure time t_(ee) represents the heating phasecontribution to the thermal exposure time t_(exp), i.e.t_(ee)=t_(peak)−t_(i) with T(t_(peak))=T_(max) (cf. also the lower partof FIG. 3 ). Similarly, the cooling exposure time t_(ce) represents thecooling phase contribution to the thermal exposure time t_(exp) and canbe defined as t_(ce)=t₂−t_(peak). It should be noted thatt_(exp)=t_(ee)+t_(ce). As can be concluded from the temperature curveaccording to FIG. 3 , the duration of the cooling phase has aconsiderable influence on the thermal exposure time t_(exp).

FIG. 4 illustrates the concept of the energy delivery time t_(ed),wherein the upper part of FIG. 4 shows examples of a right shifted (RSP)and a left shifted (LSP) normalized laser intensity pulse withI′(t)=I(t)/I_(max), wherein I_(max) is the intensity's maximal valueduring a pulse. Both the RSP pulse and the LSP pulse have an FWHM pulseduration of t_(FWHM)≈15% t_(p).

The corresponding normalized cumulative energies E′(t)=E(t)/E_(o) andnormalized temporal evolutions of the temperature increaseΔT′(t)=T(t)/ΔT_(max) are shown in the lower part of FIG. 4 . As can beseen from the Figure, the energy delivery time is much larger for theLSP pulse (t_(ed)≈88% t_(p)) than for the RSP pulse (t_(ed)≈12% t_(p)).This is understandable, since the time when the first half of the pulseenergy has been delivered to the tissue is much shorter for the LSPpulse than for the RSP pulse.

Finally, we would like to show that the energy delivery time t_(ed) ofthe laser pulse is equal to the energy exposure time tee of the tissue:

During the laser pulse, the energy of the laser pulse rapidly flows intothe superficial layer of the tissue within the penetration depth δ, andmore slowly flows out of this layer deeper into the tissue by means ofheat conduction (i.e., thermal diffusion). It is to be noted that in theabsence of thermal diffusion, the temperature increase ΔT(t) at time tduring the heating phase would be proportional with a proportionalitycoefficient k to the above-defined cumulative energy E(t) that has beendelivered to the tissue until time t. Since the rate of heat conductioncan be approximated to be also proportional to ΔT(t), the effect of heatdiffusion is simply to reduce the proportionality coefficient to asmaller value (k′), with the resulting temperature increase ΔT(t) attime t during the heating phase being proportional to the above-definedcumulative energy E(t) that has been delivered to the tissue until timet. It is noted that both the energy delivery time t_(ed) and the energyexposure time t_(ee) start at the point where the cumulative energycurve E(t) and the temperature difference curve ΔT(t), respectively,reach its first half maximum. Further, both the energy delivery timet_(ed) and the energy exposure time t_(ee) end at time t_(p) when thepulse energy of the laser pulse has been delivered. Thus, t_(ed)=t_(ee)holds.

SUMMARY OF THE INVENTION

According to one aspect of the present invention, an apparatus fortissue regeneration is provided. The apparatus comprises means forgenerating at least one laser pulse comprising a wavelength; and meansfor directing the at least one laser pulse onto a tissue surface of ahuman or animal body, wherein the means for generating comprises controlmeans to ensure that a sum of the pulse energies of the at least onelaser pulse is selected so that the corresponding fluence on the tissuesurface heats the tissue surface up to a maximal temperature T_(max)between 70° C. and a tissue boiling temperature T_(b). Further, themeans for generating of the apparatus are adapted so that a deliverytime t_(ed) of the at least one laser pulse (during which the secondhalf of the pulse energy is delivered) is sufficiently short so that,given the wavelength and thus a corresponding penetration depth δ of theat least one laser pulse, a thermal exposure time t_(exp) of the tissuesurface is shorter than 900 microseconds. Here, the thermal exposuretime t_(exp) of the tissue surface is defined as a time interval inwhich the temperature of the tissue surface is aboveT_(o)+(T_(max)−T_(o))/2, wherein T_(o) defines the initial temperatureof the tissue surface, before the laser pulse arrives.

Contrary to the conventional wisdom, it has been discovered that thetissue can be heated up to a maximal temperature T_(max) between 70° C.and T_(b) without the occurrence of significant chemical damage to thetissue (in particular protein denaturalization), if the thermal exposuretime of the tissue is sufficiently short, i.e., the tissue is heated upto high temperatures only during a short time span. In particular, ithas been discovered that the thermal exposure time t_(exp) of the tissuesurface should be shorter than 900 microseconds, preferably smaller than600 microseconds.

Preferably, the tissue is heated up to a maximal temperature T_(max)between 90° C. and T_(b) without the occurrence of significant chemicaldamage, more preferably to a maximal temperature T_(max) between 120° C.and T_(b).

It should be noted that the energy which is necessary to heat up thetissue to the above-specified temperature range can be distributed amongmore than one laser pulse. Further, the fluence (in Joule/cm²) of thelaser pulse on the tissue surface which is delivered by the at least onelaser pulse is the relevant quantity for the achievable temperature. Thefluence on the tissue surface depends on the pulse energy of laser pulseand also on how tightly this laser pulse is focused in the lateraldirection(s). Similarly, the fluence on the tissue surface is alsoinfluenced by the distance between the apparatus which emits the laserpulse and the tissue surface.

It has been further discovered that the thermal exposure time of thetissue surface can be made shorter by making the energy delivery timet_(ed) of the at least one laser pulse shorter. It should be noted thatthe time span t_(ed) during which the second half of the pulse energy isdelivered to the tissue is the relevant time span which must be keptshort, in order to avoid tissue damage, since, during this time span,the tissue is already heated up by the first half of the pulse energy.Thus, during this time span t_(ed), the tissue has an elevatedtemperature. This is also why the shape of the laser pulse plays animportant role for the present invention.

Preferably, the energy delivery time t_(ed) of the at least one laserpulse is smaller than 600 microseconds, more preferably smaller than 300microseconds and most preferably smaller than 100 microseconds.

Another quantity which influences the thermal exposure time of thetissue surface is the penetration depth δ of the laser light within thetissue. In particular, the deeper the laser light penetrates into thetissue, the longer it takes for the tissue to cool down from the maximumtemperature T_(max).

It is noted that the wavelength of the laser determines how much lightis absorbed by the tissue material. Thus, the wavelength of the laserdetermines the penetration depth δ.

Preferably, the means for generating of the apparatus are adapted toselect the wavelength of the at least one laser pulse so that thepenetration depth δ of the at least one laser pulse is smaller than 30micrometers, preferably smaller than 10 micrometers, and most preferablyshorter than 4 micrometers.

Preferably, the means for generating are adapted to select thewavelength of the at least one laser pulse between 2.6 and 3.2micrometers or between 9.1 and 10.2 micrometers, in order to coincidewith the mid-infrared water absorption peaks of water which is the majorconstituent of human and body tissues.

Quantitatively, it has been discovered that the thermal exposure time isgiven by the expression t_(exp)=t_(ed)+(1/D) (δ+√(2Dt_(ed)))², whereinD=0.1 mm² s⁻¹ is the thermal diffusivity of the treated tissue.

According to another aspect of the present invention, the treated tissuehas a critical temperature T_(crit) which is the temperature up to whichthe tissue can be heated without the occurrence of significant damages(cf. further below for a more precise definition of T_(crit)). Thecritical temperature T_(crit) depends on the thermal exposure time ofthe tissue and hence is determined by the parameters wavelength andenergy delivery time of the delivered laser pulse.

If T_(crit)≥T_(b) the critical temperature T_(crit) will not be reachedeven if the fluence F (in J/cm²) of the laser pulse is chosen so largethat the corresponding fluence on the tissue surface is greater thantwice the ablation threshold of the tissue. The reason for not reachingthe critical temperature T_(crit) is that, at the boiling temperatureT_(b), micro-explosions of over-heated tissue water occur which lead tothe ejection of tissue material from the tissue surface. The ejection oftissue material is a cooling mechanism which prevents the further riseof the temperature even if the laser pulse delivers further energy tothe tissue.

Preferably, the pulse energy (E_(o)) of the at least one laser pulse ischosen such that the corresponding fluence on the tissue surface isbelow the ablation threshold fluence, or not significantly above theablation threshold fluence, i.e., the fluence on the tissue surface canbe up to 4 times the ablation threshold fluence F_(abl) of the tissue,more preferably up to 3 times the ablation threshold fluence and, mostpreferably, 1.5 times the ablation threshold fluence.

On the other hand, if T_(crit)<T_(b), it is important that the maximumtissue temperature stays below the critical temperature T_(crit) (for agiven thermal exposure time of the tissue surface). Thus, the means forgenerating are adapted to select the pulse energy of the at least onelaser pulse sufficiently small so that the tissue surface is heated to amaximal temperature T_(max) which is smaller than the criticaltemperature T_(crit) of the tissue.

According to another aspect of the present invention, the apparatusgenerates a plurality of laser pulses (i.e., a pulse train) which aredirected to the tissue surface. In order to avoid an overheating of thetissue surface by the plurality of laser pulses, the means forgenerating are adapted to select the serial period (t_(ser)) between twosuccessive laser pulses longer than about 10 t_(exp), preferably longerthan about 50 t_(exp), and most preferably longer than about 150t_(exp).

On the other hand, in order that the deeper lying tissues do not cooldown appreciably during the time span between two ESTART laser pulses,the serial period t_(ser) should be shorter than about 3 seconds,preferably shorter than 1 second, most preferably shorter than 0.5seconds.

Further, the number N of pulses of the pulse train should be selected sothat the total duration of the pulse train t_(DMR)=N×t_(ser) is shorterthan 30 seconds, preferably shorter than 10 seconds and, most preferablyshorter than 5 seconds.

Further, instead of generating one spot S on the tissue surface by theatleast one laser pulse, in certain preferred embodiments, the means fordirecting the at least one laser pulse are adapted so that the at leastone laser pulse generates two or more spots on the tissue surface. Eachof the at least one laser pulse may be adapted such as to generate twoor more spots on the tissue surface.

Preferably, the means for directing the at least one laser pulsecomprise a screen with holes which effects the generation of the two ormore spots on the tissue surface. Alternatively, the means for directingcomprise a lens array which effects the generation of the two or morespots. Further alternatively, the means for directing comprises adiffraction optics which uses interference effects for generating thetwo or more spots on the tissue surface. The inventors of the presentapplication have realized that the approach of using a screen, whichblocks a sizeable portion of the generated pulses, leads to heatingproblems when using relatively high fluence pulses. These can beameliorated using the latter two approaches.

Preferably, the size d of each spot and the distance x betweenneighboring spots are selected so that the spots cover 25% to 65% of thetreatment area.

Preferably, the size of the spots on the tissue surface is in the rangeof 0.3 mm≤d≤1.5 mm, more preferably in the range of 0.6 mm≤d≤1.2 mm.

According to a further aspect of the present invention, a method fortissue regeneration which uses a laser system is provided. The methodcomprises the following step: directing at least one laser pulsecomprising a wavelength onto a tissue surface of a human or animal body,wherein the energy delivery time t_(ed) of the at least one laser pulse,during which the second half of the pulse energy is delivered, is chosensufficiently short, so that, given the wavelength and thus acorresponding penetration depth δ of the at least one laser pulse, athermal exposure time t_(exp) of the tissue surface is smaller than 900microseconds. Besides, the sum of the pulse energies of the at least onelaser pulse is selected so that the corresponding fluence heats thetissue surface up to a maximal temperature T_(max) between 70° C. and atissue boiling temperature T_(b).

According to a further aspect of the present invention, an apparatus fortreating male or female urinary symptoms or male erectile dysfunction isprovided. Here, the term urinary symptoms refers to at least one of thefollowing symptoms: incontinence (i.e., involuntary leakage), dysuria(i.e., painful urination), urgency and frequency of urination, orrecurrent infections. The apparatus comprises means for generating atleast one laser pulse comprising a wavelength; and means for introducingthe at least one laser pulse into the urethra. The apparatus maycomprise control means to ensure that a sum of the pulse energies of theat least one laser pulse is selected so that the corresponding fluenceon the surface of the urethra heats the tissue surface up to a maximaltemperature T_(max) between 70° C. and a tissue boiling temperatureT_(b). Further, the means for generating of the apparatus may be adaptedso that a delivery time t_(ed) of the at least one laser pulse (duringwhich the second half of the pulse energy is delivered) is sufficientlyshort so that, given the wavelength and thus a corresponding penetrationdepth δ of the at least one laser pulse, a thermal exposure time t_(exp)of the surface of the urethra is shorter than 900 microseconds.

Here, it should be noted that the upper layer of the urethra is a mucosalayer which has a tissue boiling temperature T_(b) which isapproximately 250° C.

By introducing such laser pulses into the urethra, a mild hyperthermiais induced in the urethra and in the surrounding tissues.

When treating the male urethra, the afore-mentioned surrounding tissueconsists of the periurethral erectile tissue (corpus spongiosum) whichleads to the generation of new vessels, a regeneration of connectivesupport and an improvement of the vascular function. As the vasculartissue and the fibro-connective support in all the tissue compartmentsis responsible for the maintenance of an erection, the laser therapywhich uses the above-described apparatus helps in curing erectiledysfunction.

Similarly, when treating urinary symptoms it is known that the thicknessof the urethral mucosa and the rich vascularization of the submucosaconfer its sealing properties, since it is the main contributory factorof the reduction of both the caliber of the urethral opening and theradius of the urethral cylinder. The superficial warming process inducedin the urethra improves the urethral trophism by promoting thevasodilation effect, and thus improves the performance of the intrinsicmechanism of continence by decreasing the radius of the urethra due tothe improvement in the submucosal vascular plexus and the improvement inthe thickness of the epithelium.

Preferably, the energy delivery time t_(ed) of the at least one laserpulse is smaller than 600 microseconds, more preferably smaller than 300microseconds and most preferably smaller than 100 microseconds.

Preferably, the means for generating of the apparatus are adapted toselect the wavelength of the at least one laser pulse so that thepenetration depth δ of the at least one laser pulse is smaller than 30micrometers, preferably smaller than 10 micrometers, and most preferablyshorter than 4 micrometers.

Preferably, the means for generating of the apparatus comprise two lasersystems, wherein the two laser systems operate at different wavelengths.

The advantage of using two different wavelengths is that the twodifferent wavelengths can achieve different clinical effects.

Preferably, the means for introducing the at least one laser pulse intothe urethra comprises a cannula; and a handpiece which guides the atleast one laser pulse to a treatment area on the surface of the urethra.

Preferably, the handpiece is fixedly connected to the cannula during thetreatment of the urethra.

Preferably, the apparatus further comprises means for inspecting theurethra.

Preferably, the means for inspecting the urethra comprises a cannula;and an endoscope.

A further aspect of the present invention relates to the use of one ormore laser pulses for treating urinary symptoms or erectile dysfunction.

According to a further aspect of the present invention, a method fortreating urinary symptoms or erectile dysfunction is provided. Themethod comprises the following steps: introducing at least one laserpulse comprising a wavelength into the urethra; and guiding the at leastone laser pulse to a treatment area on the surface of urethra, whereinthe energy delivery time t_(ed) of the at least one laser pulse, duringwhich the second half of the pulse energy is delivered, is chosensufficiently short, so that, given the wavelength and thus acorresponding penetration depth δ of the at least one laser pulse, athermal exposure time t_(exp) of the tissue surface is smaller than 900microseconds. Besides, the sum of the pulse energies of the at least onelaser pulse is selected so that the corresponding fluence heats thetissue surface up to a maximal temperature T_(max) between 70° C. and atissue boiling temperature T_(b).

It is noted that the present invention encompasses all combinations ofthe above-described features, unless such a combination is not feasible(as it is, for example, self-contradictory or not working). Inparticular, the features described above with reference to an apparatusmay be implemented as corresponding method steps. Also, the aspects andsteps described in the present application may specifically be appliedalso for treating the urethra, even if not expressly mentioned.

The conventional approach for tissue regeneration focuses on a slowthermal pulsing of the connective tissue, in order to stimulatefibroblasts and other cells to respond to wound healing scenarios. Incontrast, the approach according to the present invention is based on anindirect mechanism whereby the keratinocytes and other cells located inthe superficial layer above the basement membrane are activated (i.e.,triggered) by extremely fast and intense “heat shock” thermal pulses.The new collagen production is thus triggered not by the directtemperature elevation of the connective tissue but instead by the crosstalk between the deeper lying fibroblasts and the superficially locatedkeratinocytes that are activated by the fast thermal pulse triggeringdelivered by our innovative apparatus and method. In what follows theterm ESTART (Epithelium Superficially Triggered Activation ofRegeneration of Tissue) will be used from time to time to describe thisnew treatment apparatus and method.

The rapidly heated thin superficial tissue layer is then quickly cooleddown by the fast diffusion of the generated heat to the underlyingcolder tissue layers. This is caused by the temperature gradient, whichis very large in the superficial thin layer. As a result, the conductivecooling of the superficial layer is very effective, in contrast towithin deeper connective tissue layers that remain moderately heated fora significantly longer time, owing to weaker heat flow from this region.Since the protein denaturation rate depends on temperature in a highlynonlinear manner, the short-lived high temperature gradient can be usedto trigger deeper tissue regeneration without thermally damaging theepithelial tissue.

Therefore, as opposed to connective tissue heating where heat diffusionprolongs exposure times, it is the fast thermal diffusion from theheated thin superficial tissue layer that according to our innovationfacilitates the generation of extremely short thermal exposure pulses.

The approach according to the present invention resembles themicro-needling technique that aims not to injure the keratinocytes butto stimulate them with superficial punctures and without any aggressionto fibroblasts. Micro-needling has been introduced as a significantlyless aggressive alternative to skin resurfacing or dermabrasion. Skinresurfacing and dermabrasion as used in aesthetic medicine for improvingskin quality are based on “ablation” (destruction or wounding ofsuperficial skin layers), which requires several weeks for healing thatinvolves formation of new skin layers. Such procedures provoke an acuteinflammatory response. A much less intense inflammatory response occursfollowing a microneedle perforation of the skin. The mechanism of actionof micro-needling appears to be different from the mere inflammatoryresponse to the localized “ablation”, and seems to involve also inducedcell proliferation by electrical signals. Since with micro-needling onlya relatively small percentage of the skin is being affected, thetreatment outcome is expected to be accordingly limited. On the otherhand, the innovative ESTART concept according to our innovation can beviewed as a non-ablative thermal “needling” (i.e., triggering) of thetotal treated skin surface, with the action of the spatially sharpneedles being replaced by the action of temporarily “sharp” butspatially broad thermal pulses.

The ESTART concept also resembles the laser induced thermalpre-conditioning. In this technique, the skin is pre-conditioned bybeing submitted to mild laser-induced heat shock prior to surgery.Laser-preconditioned incisions have been found to be two times strongerthan control wounds that had not been laser pre-treated. However, thelaser induced thermal pre-conditioning is performed over very longexposure times, on the order of minutes and therefore at temperaturesbelow 50° C., and does not involve only epithelial tissues.

The ESTART concept may have applications in skin rejuvenation,gynecology, urology, dentistry and other medical areas. For example, indentistry, ESTART may be used to promote wound healing in gingivaltissue by increasing the number of fibroblasts. Additionally,regenerative effects on the alveolar bone are also expected.

DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS

Some of the embodiments of the invention will be explained in thefollowing with the aid of the Figures in more detail. It is shown in

FIG. 1 a schematic illustration of the tissues surrounding the internaland external body surfaces;

FIG. 2 an illustration of the penetration depth δ within the irradiatedtissue;

FIG. 3 the intensity of the delivered laser pulse (above) and theresulting thermal pulse on the tissue surface (bellow);

FIG. 4 the intensity of a of a “right shifted” and a “left shifted”laser pulse with the same FWHM energy pulse duration (above) andcorresponding cumulative energy and thermal curves (below);

FIG. 5 an apparatus for tissue generation according to the presentinvention which comprises a laser system;

FIG. 6 the curve of critical temperature T_(crit) over the thermalexposure time using standard Arrhenius parameters;

FIG. 7 the critical temperature over the thermal exposure times forhuman tissue which is based on clinical observations;

FIG. 8 the surface temperature T_(max) at the end of a laser pulse as afunction of the normalized fluence F/F_(abl);

FIG. 9 the dependence of the thermal exposure time t_(exp) on the energypenetration depth δ for different energy delivery times t_(ed);

FIG. 10 the serial delivery of laser pulses with a goal to achieve dualmechanism regeneration;

FIG. 11 the temperature build-up during the serial delivery of laserpulses;

FIG. 12 the spatially patterned delivery of laser pulses;

FIG. 13 applicator elements for a laser treatment of the male or femaleurethra

a) Exposure of the Tissue to a Laser Pulse

An embodiment of the apparatus for tissue regeneration according to thepresent invention is the laser system 1 shown in FIG. 5 . This lasersystem includes a laser device 2, a laser handpiece 5 and a control unit4. The laser device generates a laser beam, while control unit is usedfor controlling and modifying the laser beam. Thus, the laser device andthe control unit taken together are an embodiment of the means forgenerating according to the present invention.

Further, the laser handpiece 5 emits the laser beam so that it impingeson the tissue surface. Thus, the laser handpiece is an embodiment of themeans for directing according to the present invention. Alternatively,an optical imaging system 6 (for example, a lens system) can be arrangedbetween the laser handpiece and the tissue surface. Such an imagingsystem modifies the laser beam being emitted from the laser handpieceand then directs the modified beam to the tissue surface.

On the tissue surface, the extension of the laser beam is given by thespot S, wherein the spot S can be defined as the region of the tissuesurface within which 90% of the total energy of the laser pulse isdelivered to the tissue (i.e., the spot S corresponds to the smallestregion on the tissue surface so that, during the delivery of the laserpulse, 90% of the total energy of the laser pulse is delivered to thisregion), cf. also FIG. 5 . The spot S can also be referred as thetreatment area. The laser system according to FIG. 2 operates in pulses,i.e. at least one laser pulse is emitted from the handpiece. Typically,the laser system generates a plurality of laser pulse, i.e. a train oflaser pulses which are directed to the tissue surface.

As mentioned above, the laser pulse heats up the surface of theirradiated tissue to a maximal temperature T_(max), wherein the heatingphase lasts until the time when the pulse energy of the laser pulse hasbeen delivered. Thereafter, the cooling phase of the tissue begins (cf.also the above definitions of the energy exposure time t_(ee) and thecooling exposure time t_(ce)).

It is noted that the rate of the cooling can be enhanced by applying anexternal cooling to the epithelial surface before or immediately afterthe energy delivery. External cooling may be accomplished for example byusing a cryogenic or liquid (for example, water) spray, a contactcooling or forced air.

b) Tissue Damage (Arrhenius Integral)

In the following, tissue damage means chemical damage of the tissue, inparticular protein denaturalization which is an irreversible chemicalreaction. It is noted that protein denaturalization is the most relevantprocess of all the chemical processes which lead to tissue damage.

Commonly, the metric for tissue damage (Ω) is the ratio of theconcentration of native (undamaged) tissue before thermal exposure(C_(o)) to the concentration of native tissue at the end of the exposuretime (C_(f)). The equation for the Arrhenius integral isΩ=ln(C _(o) /C _(f))=A∫ exp(−E/RT)dt  (2)where Ω is the tissue damage, A is the frequency factor (i.e. the damagerate in 1/sec), E is the activation energy (in J/mol), T is thetemperature of exposure (in degrees K), R is the gas constant (R=8.32J/mol K) and the integral is over the duration of the thermal pulse, Δt.Here, a square-shaped thermal pulse with a constant temperature duringthe duration of the thermal pulse is assumed.

The tissue damage kinetics can be characterized by a criticaltemperature (T_(crit)) which is defined asT _(crit) =E/(Rln(AΔt))  (3)and represents the temperature at which the concentration of theundamaged tissue is reduced by a factor of e (i.e., Ω=1). While for longexposure times it is relatively easy to achieve approximatelysquare-shaped thermal pulses, this becomes exceedingly difficult forexposure times shorter than approximately 1 second where thecontribution of the slowly falling temperature during the cooling phaseto the thermal exposure time becomes appreciable. Thus, for shortexposure times, the shape of the temperature pulse more resembles aquasi “triangular” shape as represented in FIG. 3 above. For this reasonand in order to be able to calculate critical temperatures for extremelyshort pulse durations, we redefine the Arrhenius parameters in such amanner that the critical temperature represents the maximal temperatureduring a FWHM thermal exposure time t_(exp) for which Ω=1. Note thatthis does not affect the calculation of the critical temperature forlonger pulse durations where nearly square shaped thermal pulses can beachieved, and therefore t_(exp)≈Δt.

Published Arrhenius parameters vary but typical values (the “standard”values) for the Arrhenius parameters for skin which are obtainedexperimentally for exposure times t_(exp)≥1 sec are as follows: A=3.010⁸⁷ s⁻¹ and E=5.5 10⁸ J/kmol. FIG. 6 shows the calculated expecteddependence of the critical temperature on the exposure time, based onthese standard parameters.

During standard regeneration treatments, the deeper lying layers ofconnective tissue are the tissue layers which one tries to thermallyinjure directly. Since deeper tissue layers remain heated for a longtime (because of the weak flow of heat to the surrounding unheatedtissues), this results in long cooling exposure times (t_(ce)), andconsequently long overall exposure times (t_(exp)) of the connectivetissue. Typical exposure times for the connective tissue are in theorder of 10 seconds or longer. Consequently, as seen in FIG. 6 ,critical temperatures for the connective tissue are in the relativelylow temperature range of 50 to 70° C. It is noted that, according to thestandard Arrhenius parameters, the critical temperatures for theepithelium are similarly low.

During our clinical tests, however, we made the surprising discoverythat, at extremely short exposure times, the viability of the skin cellsis much higher than what would be expected from the observed cellviability for long thermal pulse durations (i.e., from the standardArrhenius parameters). A dramatic change in the slope of the criticaltemperature versus thermal exposure time was observed (see our measureddata on human skin represented by circles in FIG. 6 ). The exact reasonfor this change is not completely understood, but could be related tothe cellular repair mechanism taking place at multi-second durations.The deviation from the Arrhenius law for shorter exposure times can notonly be inferred from our limited clinical observations but also fromthe “Standard Guide for Heated System Surface Conditions that ProduceContact Burn Injuries”, as published by The American Society for Testingand Materials (ASTM). This standard guide includes a chart which relatescontact skin temperature to the time needed to cause a burn. Thecritical temperature curve is shown to start deviating from the standardArrhenius law for thermal exposure times which are shorter than 10 sec.In particular, the critical temperatures for thermal exposure timest_(exp)≤1 sec are higher than what would be expected from the standardArrhenius relationship.

Taking into account our clinically observed critical temperatures atshort exposure times (represented by circles in FIGS. 6 and 7 ), wedeveloped a prediction model for the critical temperatures over thecomplete range of exposure times (full line in FIG. 7 ). Our model isbased on a discovery that the cell viability of any type of tissue canbe described as a combined effect of two biochemical processes, P₁ andP₂, that dominate cell survival characteristics at very long (processP₁) and very short (process P₂) exposure times (See FIG. 7 ). Ouranalysis shows that the short exposure time process P₂ is characterizedapproximately by A=8.0 10⁴ sec⁻¹ and E=1.8 10⁷ J/kmol, while the longexposure time process (P₂) is characterized by the standard Arrheniusparameters as described above. It is to be expected that when attemptingthe regeneration of different types of tissues, critical temperatureswill vary to a certain degree from tissue to tissue. However, since skinis known to be highly sensitive to temperature, it can be taken thatFIG. 7 represents the worst case, and that for other types of epitheliumand connective tissue, the critical temperatures will only be higher butnot lower than what is shown in FIG. 7 . Therefore, the safetyconsiderations as described in this invention apply to all types ofepithelium and connective tissues.

It is to be noted that ESTART treatments are based primarily on thecharacteristics of the short exposure time-biochemical process P₂.According to the curves shown in FIG. 7 , no irreversible thermal injuryis expected for temperatures up to and even above 250° C. provided thatthe thermal exposure time is shorter than approximately 1 msec.

c) Confined Water Boiling

In regeneration treatments, the goal is to locally heat up the tissue(in contrast to selective thermolysis treatments such as hair removal orvascular treatments where the goal is to selectively heat a specificspatially localized chromophore within the tissue). Since water is themajor constituent of biological tissue, this means that the goal ofregeneration treatments is to homogeneously heat up the tissue water. Inthe present method for tissue regeneration, the tissue water is heatedup directly by absorbing the delivered laser light.

For example, if the epithelia is irradiated by radiation from an Er:YAGlaser with wavelength λ=2,940 nm, the radiation is strongly absorbedwithin the superficial layer with a thickness of δ≈1 μm. Gas CO₂ lasersemit at several wavelengths within the water absorption peak at 9 to 12μm. The penetration depth in water of the CO₂ laser with wavelengthλ=10,640 nm is approximately 20 times larger than for Er:YAG.

A microscopic numerical physical model was developed for the tissuewater which is heated up directly within the penetration depth (δ),wherein the penetration depth represents the inverse of the absorptioncoefficient within the tissue at the wavelength of the delivered laserpulse. Assuming that water is the major absorber, the tissue was treatedas a homogenous material throughout epithelia and connective tissue. Thethermodynamic behaviour of tissue water was combined with the elasticresponse of the surrounding solid tissue medium. This was complementedby a one-dimensional treatment of heat diffusion using afinite-difference scheme and by modelling protein denaturation kineticswith the Arrhenius integral.

It is to be appreciated that the tissue water plays an important role inour invention, since the heating of the tissue water to the confinedboiling temperature (T_(b)) results in micro-explosions which ejectover-heated tissue from the tissue surface. Thus, confined boiling ofthe tissue water leads to tissue ablation.

It should be further noted that this ablation mechanism which is basedon confined boiling is different from other ablation mechanisms whichinvolve strong acoustic transients, plasma formation, or transientbubble formation and which occur at higher laser intensities.

Further, it is noted that the ejection of over-heated tissue from thetreated tissue acts as a cooling mechanism for the remaining tissue. Asa consequence, even if further laser light is absorbed by the tissuewater, the temperature of the treated tissue cannot increase beyond thetemperature T_(b) where confined boiling occurs, i.e., the maximalsurface tissue temperature is effectively kept at T_(max)=T_(b). FIG. 8shows the maximal surface temperature T_(max) of the tissue as afunction of the normalized laser fluence F/F_(abl).

Our measurements and calculations show that the ablation thresholdfluence F_(abl) can be calculated from the known penetration depth δ as:F _(abl) =H _(a)×δ  (4)wherein H_(a) is the specific heat of ablation for soft tissue withH_(a)≈1.5×10⁴ J/cm³. As an example, F_(abl)=1.5 J/cm² for the Er:YAGlaser.

The measured and calculated values of T_(b) for human soft tissues varyto a certain degree (i.e., 248° C.≤T_(b)≤258° C.). Our calculations fora soft tissue give T_(b)=256° C. For simplicity, we shall further assumein the following that T_(b)≈250° C.

Therefore, as long as the thermal exposure time t_(exp) is kept so shortthat the corresponding critical temperature T_(crit) of the tissue isgreater or equal to the confined boiling temperature (T_(crit)≥T_(b)),the confined boiling of the tissue water will act as a regulatingmechanism limiting which keeps the tissue temperature away from thecritical temperature T_(crit). As heating the tissue to the criticaltemperature T_(crit) involves irreversible chemical tissue damage(protein denaturalization), the confined water boiling (CWB) mechanismrepresents a significant safety feature.

This safety mechanism is helpful in many practical circumstances. Forexample, when treating body cavities, the internal surface may not beregularly shaped and hence the delivered fluence, F (in J/cm²) and thegenerated heat may vary from one location to another. Besides, theenergy source may not be always kept at the optimal distance or optimalangle with regard to the tissue surface, again resulting in anon-uniform heat pulse generation. And finally, a physician error canalso occur. In all these cases, the CWB mechanism protects the patientfrom any irreversible injury and keeps the treatment within the safeESTART limits.

It should be noted that the above-described ablation mechanism which iscaused by the confined boiling of the tissue water is not harmful to thetreated tissue. This is, since only a very thin layer of the tissue isremoved by this ablation mechanism, providing that the fluence is notsignificantly above the ablation threshold. In particular, a typicalablation rate with an Er:YAG laser is 4 microns/(J/cm²). Therefore, evenif the fluence exceeds the ablation threshold of about 1.5 J/cm² by afluence factor k_(f):=F_(o)/F_(abl) in the range of k_(f)=1.25 to 4, themaximal thickness of the removed layer does not exceed 3 to 18 microns.In the case that skin is treated, the removed layer is even smaller thanthe layer of dead cells on the surface of the skin which is constantlyshredded away, in order to make room for newer cells to replace the oldcells.

It should also be noted that our inventive regeneration apparatus andmethod are primarily intended to be non-ablative. However, in case thedelivered fluence exceeds by accident or by intent or for any otherreason the ablation threshold, the ablation by means of confined waterboiling will not be harmful for the tissue, but will prevent harmfultissue damage due to protein denaturalization (if T_(crit)≥T_(b)). Thisfeature allows the practitioner to set the nominal fluence to a levelwhere temperatures up to T_(b) can be expected, without the risk oftissue damage in case the actual fluence during the treatment starts todeviate from the set nominal fluence.

d) Dependence of the Thermal Exposure Time t_(exp) on the PenetrationDepth δ and the Energy Delivery Time t_(ed)

In what follows, it will be assumed that the fluence F is homogeneousover the spot S. In case that the fluence is not homogeneous, themaximal fluence within the reduced spot S′ over which the maximalfluence may be considered approximately homogeneous is the quantity tobe considered. Also, it is the assumption that the dimensions of thespot S are such that the thermal diffusion in the lateral directionduring the thermal exposure time can be ignored. Here, the lateraldirection is to be understood as the direction along the tissue surfaceas opposed to the vertical direction that is directed into the tissue.When considering a circular spot S with a diameter d the thermalrelaxation time (TRT) in the lateral direction can be calculated fromthe relation TRT=d²/16 D where D=0.1 mm²/s is the thermal diffusivity ofthe soft tissue. As will be shown below, thermal exposure times shorterthan about 5 ms are of particular interest for the present innovation.It can be therefore concluded that for spot diameters larger than aboutd=0.1 mm (TRT≈5 ms), the thermal diffusion in the lateral direction doesnot have a significant influence on t_(exp). It should be appreciatedthat the spot S does not have to be circular. Therefore, the parameter dshould be understood as the smallest dimension of the spot S in thelateral direction.

It is important to realize that the thermal exposure time (t_(exp)) isdetermined by the two quantities, penetration depth (δ) and energydelivery time t_(ed). This can be understood as follows: as shown above,t_(exp)=t_(ed)+t_(ce) holds. Further, the cooling exposure time t_(ce)is determined by the penetration depth δ, since, as mentioned above, thecooling time depends on the thickness of the layer which has been heatedup by the laser pulse. Our analysis shows that the cooling exposure timedepends on δ and t_(ed) as

$\begin{matrix}{t_{ce} \approx {\frac{1}{D}\left( {\delta + \sqrt{2{Dt}_{ed}}} \right)^{2}}} & (5)\end{matrix}$

The physics behind this equation is as follows: the cooling time t_(ce)of the superficial tissue layer after the heating phase can becalculated from t_(ce)=(1/D) d_(h) ², where d_(h) represents thediffusion length d_(h) of the heated superficial tissue layer at the endof the heating phase and D≈0.1 mm² s⁻¹ is the thermal diffusivity of thetissue. The thickness d_(h) is equal to the penetration depth δ plus anadditional diffusion distance d_(d) resulting from heat diffusion duringthe heating phase, i.e., d_(h)=δ+d_(d). It is known that the diffusiondistance following a certain diffusion time t_(d) is proportional to √(Dt_(d)), wherein the exact relation depends on the actual geometrical andtemporal conditions. By fitting this expression to the numerical model,we determined that, for the conditions considered in the presentinvention, the diffusion distance is best described by d_(d)≈√(2 Dt_(ed)).

This dependence of the thermal exposure time t_(exp)=t_(ed)+t_(ce) onthe penetration depth δ and the energy delivery time t_(ed) isillustrated in FIG. 9 . As can be seen in FIG. 9 which is concerned withtypical tissue regeneration lasers such as Er:YAG or CO₂ laser, theexposure time is t_(exp)≈0.7 ms for the Er:YAG laser (λ=2,940 nm, δ≈1μm) and t_(exp)≈4.6 ms for the CO₂ laser (λ=10,640 nm, δ≈15 μm), whereinthe energy delivery time of both lasers is t_(ed)=0.175 ms. Looking nowat FIG. 7 , the critical tissue temperature for the treatment with thisparticular Er:YAG device is T_(crit)≈260° C., while the critical tissuetemperature for the treatment with this particular CO₂ laser device ismuch lower, namely T_(crit)≈105° C.

Since the critical temperature T_(crit) of the treated tissue depends onthe thermal exposure time t_(exp) (cf. the above-discussed Arrheniuscurves) and since the thermal exposure time t_(exp) depends on thepenetration depth δ and the energy delivery time t_(ed), it follows thatthe critical temperature T_(crit) of the treated tissue is alsodetermined by the two parameters of the laser device, namely thepenetration depth δ of the laser beam and the energy delivery timet_(ed) of the laser pulse.

As shown in FIG. 7 , the critical temperatures have the values ofT_(crit)=120° C., 180° C. and 250° C. for the corresponding thermalexposure times t_(exp)=3.5 ms, 2 ms and 0.8 ms. Therefore, using FIG. 9, the maximal tissue surface temperatures T_(max)=120° C., 180° C. and250° C. which can be safely imposed (i.e., T_(max)≤T_(crit)) are subjectto the condition that the corresponding optical penetration depth valuesof the treatment device are smaller or equal to about δ=18 μm,11 μm and9 μm, wherein these values of δ are the maximal allowed opticalpenetration depths for the theoretical limit of infinitely short laserpulses. Similarly, using FIG. 9 again, for infinitely short penetrationdepths δ, the maximal superficial temperatures above T_(max)=120° C.,180° C. or 250° C. can be safely imposed, if the energy delivery timest_(ed) are shorter than correspondingly about t_(ed)=1.2 ms, 0.7 ms and0.25 ms. With larger penetration depths, these limiting energy deliverytimes are correspondingly shortened.

e) Relevant Laser Parameters

As noted before, the penetration depth δ is the inverse of theabsorption coefficient of the tissue at the wavelength λ of the laserbeam. Thus, for a given tissue to be treated, the wavelength of thelaser beam determines the penetration depth δ.

Therefore, by choosing appropriate laser parameters, namely anappropriate wavelength and an appropriate intensity pulse duration andshape, the thermal exposure time t_(exp) (which, as seen above, isdetermined by the parameters δ and t_(ed)) and hence the correspondingcritical temperature T_(crit) for the treated tissue can be determined.

Using the relations shown in FIG. 9 for typical skin resurfacing laserssuch as Er:YAG or CO₂, the thermal exposure time is t_(exp)≈0.4 ms forthe Er:YAG laser (λ=2,940 nm, δ≈1 μm) and t_(exp)≈3.9 ms for the CO₂laser (λ=10640 nm, δ15 μm). These numbers assume that the energydelivery times for both lasers of t_(ed)=0.1 ms. Taking into accountFIG. 7 , this translates into T_(crit)>300° C. for the Er:YAG laser andT_(crit)=110° C. for the CO₂ laser. Similarly, the Er,Cr:YSGG laser(λ=2,780 nm, δ=3 μm) with t_(ed)=0.1 ms, has an exposure time oft_(exp)≈0.7 ms which translates into T_(crit)≈250° C. It follows that,for this energy delivery time, only the Erbium lasers have a criticaltemperature above or at the boiling temperature T_(b)=250° C.

Laser parameters which lead to a critical temperature above the boilingtemperature (T_(crit)≥T_(b)) are particularly attractive for being usedby the present apparatus and method for tissue regeneration, since, inthis case, the tissue regeneration treatment can be safely operatedwithout a special control of the fluence (F) of the laser pulse. Thisis, since the above-discussed confined water boiling (CWB) ensures thatthe critical temperature is not reached. Thus, the fluence of the usedlaser pulse can be several times above the ablation threshold.

Our analysis (see Table 1) shows that such a safe operation regime fortissue regeneration (T_(crit)≥T_(b)) can be achieved with laserparameters which lead to a penetration depth of less than or equal to 6μm provided that the energy delivery time is less than or equal to 50μs. Similarly, for penetration depths less than or equal to 4 μm, theenergy exposure time should be less than or equal to 100 μs, preferablyless than or equal to 80 μs. For penetration depths less than or equalto 1 μm, the energy delivery time should be less than or equal to 250μs, preferably less than or equal to 200 μs.

In order to achieve an extremely short penetration depth, the lasershould have a wavelength which is close to the peak absorption for waterat λ≈3,000 nm. Examples of such devices are the Er:YAG laser withλ=2,940 nm (with a corresponding penetration depth in water of 1 μm), orthe Er,Cr:YSGG laser with λ=2,780 nm (penetration depth in water of δ≈3μm). This leads to the requirement that the energy delivery time fromthe Er:YAG laser is chosen sufficiently short that the energy deliverytime is shorter than or equal than about 250 μs, preferably less than orequal to about 200 μs. For the Er,Cr:YSGG laser, the energy deliverytime of the laser pulse should be chosen sufficiently short that theenergy delivery time is shorter than or equal to about 150 μs,preferably less than or equal to about 100 μs.

Laser pulses which lead to a critical temperature below the boilingtemperature (T_(crit)≤T_(b)) can also be used by the present apparatusand method for tissue regeneration provided that the fluence (F) of theenergy pulse within the spot S is controlled such that the maximaltemperature T_(max) does not exceed the critical temperature of thetissue. In contrast, such a control of the fluence is not required ifT_(crit)≥T_(b), as the CWB mechanism ensures that T_(max) remains belowor at T_(crit).

For T_(crit)≤180° C. and for a penetration depth less than or equal to10 μm, the energy exposure time should be less than or equal to 100 μs,preferably less than or equal to 50 μs. Similarly, for the device withthe penetration depth less than or equal to 4 μm the energy deliverytime should be less than or equal to 350 μs, preferably less than orequal to 300 μs. And for the penetration depth less than or equal to 1μm the energy delivery time should be less than or equal to 600 μs,preferably less than or equal to 500 μs.

For T_(crit)≤120° C., and with the penetration depth less than or equalto 15 μm, the energy delivery time should be less than or equal to 100μs, preferably less than or equal to 50 μs. Similarly, with thepenetration depth less than or equal to 10 μm, the energy delivery timeshould be less than or equal to 350 μs, preferably less than or equal to300 μs. And, for the penetration depth less than or equal to 4 μm theenergy delivery time should be less than or equal to 800 μs, preferablyless than or equal to 700 μs.

For T_(crit)≤80° C., and with the penetration depth less than or equalto 30 μm, the energy delivery time should be less than or equal to 1 ms,preferably less than or equal to 0.8 ms. Similarly, with the penetrationdepth less than or equal to 15 μm, the energy delivery time should beless than or equal to 3.5 ms, preferably less than or equal to 3 ms.And, for the penetration depth less than or equal to 4 μm the energydelivery time should be less than or equal to 6 ms, preferably less thanor equal to 5.5 ms.

TABLE 1 T_(crit) δ t_(ed) 250° C. ≤6 μm ≤50 μs ≤4 μm ≤80-100 μs ≤1 μm≤200-250 μs 180° C. ≤10 μm ≤50-100 μs ≤4 μm ≤300-350 μs ≤1 μm ≤500-600μs 120° C. ≤15 μm ≤50-100 μs ≤10 μm ≤300-350 μs ≤4 μm ≤700-800 μs  80°C. ≤30 μm ≤0.8-1 ms ≤15 μm ≤3-3.5 ms ≤4 μm ≤5.5-6 ms

The single pulse fluences F required to elevate the surface temperatureup to the maximal temperature T_(max) can be calculated using FIG. 8 andEq. 4 asF(T _(max))=H _(a)δ(T _(max) −T _(o))/(T _(b) −T _(o));T _(max) ≤T_(b)  (6)

Assuming an Er:YAG laser (δ≈1 μm) and an initial tissue temperature ofT₀=₃₅° C., and taking into account the slight dependence of thethreshold fluence on energy delivery time (threshold is higher forlonger t_(ed)), the required fluences for exemplary T_(max) are as shownin Table 2.

TABLE 2 T_(max) (° C.) F (J/cm²) 80 0.2-0.4 120 0.5-0.7 180 0.9-1.1 2561.4-1.6

It should be noted that our innovation is based on a linear interactionof the delivered laser pulse with the tissue. Therefore, whenconsidering the reduction of the energy delivery time t_(ed) (i.e., theduration of energy delivery) to the nanosecond or even shorter range,care must be taken that the delivered peak power of the laser pulse doesnot exceed the threshold for the non-linear interaction, since, in thiscase, the tissue would get damaged due to the tissue ionization at veryhigh temperatures. Therefore, the energy delivery time t_(ed) should notbe shorter than approximately 100 ns.

f) Dual Mechanism Regeneration (DMR) Treatment

It is noted that a laser pulse is considered an ESTART laser pulse whenits optical penetration depth δ and energy delivery time t_(ed) are suchthat the critical temperature (T_(crit)) of the treated tissue (asdefined by FIG. 7 , and the corresponding Arrhenius parameters) is above70° C., preferably above 120° C., even more preferably above 180° C.,and most preferably above 250° C. Further, an ESTART* pulse is an ESTARTpulse, characterized by T_(crit)≥250° C.

A particular embodiment of our invention is the dual mechanismregeneration (DMR) apparatus which combines the fast (ESTART)superficial triggering of the epithelial surface primarily involving theshort exposure process P₂ with the conventional temperature elevation ofthe deeper lying tissues primarily involving the long exposure processP₁. This is accomplished by generating and delivering a finite series ofN ESTART laser pulses to the epithelial tissue, wherein the individuallaser pulses are temporally separated from each other with a serialperiod t_(ser) which is longer than the time required for the tissuesurface to appreciably cool down after the preceding ESTART laser pulse,and shorter than 5 s, preferably shorter than 2 s, and most preferablyshorter than 0.5 s, in order for the deeper lying (bulk) tissue not tocool down appreciably in-between ESTART pulses. In such a case thetissue temperature slowly builds up during the serial delivery of ESTARTpulses, and the long exposure process P₁takes place. In what follows weshall denote such a serial delivery of ESTART laser pulses, whichresults in a combined short exposure/long exposure treatment as a DMRpulse delivery and treatment.

FIG. 10 shows a typical temporal evolution of the tissue surfacetemperature T during a repetitive ESTART irradiation which, as anexample, uses N=6 ESTART laser pulses with a serial period of t_(ser)=50ms.

We characterize this DMR temporal evolution by two types of peaktemperatures: i) “fast” temperature peaks T_(p2i), belonging toindividual ESTART laser pulses i (with i=1 to N) within the sequence ofN pulses; and ii) a “slow” temperature peak T_(p1) which describes theraise of the “base-line” temperature of the tissue surface from thefirst to the last ESTART pulse, before the temperature of the tissuesurface decays again in the absence of further energy delivery. It isnoted all the temperatures which are shown in FIG. 10 are measured atthe centre of the spot S.

Roughly speaking, the tissue temperature T_(p1) and the correspondingexposure time t_(exp1) can be considered to be determined primarily bythe long exposure process P₁, while the tissue temperature T_(p2i) andcorresponding t_(exp1) is determined primarily by the characteristics ofthe short exposure process P₂. As described earlier, using a laserconforming to the ESTART conditions, the amount of cell injury that isincurred during fast temperature pulsing is limited, due to the veryhigh critical temperatures under such short thermal exposure time(t_(exp2)) conditions. On the other hand, the tissue injury caused bythe thermal base-line “pulse” (which lasts for seconds) can beconsiderable already for temperatures T_(p1) above 55-65° C., as can beseen from FIG. 7 . Therefore, controlling the slowly varying temperatureelevations at the surface (and deeper within the tissue) during the DMRdelivery is of critical importance for assuring the safety of theprocedure.

The specific heat capacity of water which is the major constituent oftissues is C_(h)=4.2 J/cm³ K. The cumulative fluence F_(c)=N×F_(o) thatis required to heat up the tissue volume extending to a depth z_(h) ofthe tissue for an average temperature increase of ΔT_(h) can thus beestimated as F_(c)≈z_(h)C_(h)ΔT_(h). Taking ΔT_(h)≈30° C., we obtainF_(c)≈2.5 J/cm² for z_(h)≈0.2 mm, F_(c)≈13 J/cm² for z_(h)=1 mm, andF_(c)≈130 J/cm² for z_(h)=10 mm. Therefore, the depth z_(h) of thethermally treated tissues according to the slow exposure process P₁ isdetermined by the cumulatively delivered fluence F_(c)=N×F_(o) over thesame spot S during a DMR pulse train (as set on the laser system'scontrol box). In order not to “overheat” the deeper lying tissues, thecumulative fluence should not exceed F_(c)≈150 J/cm², preferably shouldbe lower than 50 J/cm², and most preferably should be lower than 15J/cm².

Similarly, the maximal surface temperature T_(p1) is determined bysetting the serial period t_(ser) and the number N of the deliveredESTART pulses at control box. Our measurements and calculations showthat, for a laser spot dimension of d>1.5 mm, the time span during whichthe maximal temperature difference ΔT_(p21)=T_(p21)−T_(o) of the firstESTART pulse drops down to ΔT_(f1)=kΔT_(p21), with k=0.01, k=0.02 andk=0.03 depends on t_(exp) as t_(1%)≈150 t_(exp), t_(2%)≈50 t_(exp), andt_(3%)≈10 t_(exp), correspondingly. The goal of the DMR serial deliveryof pulses is that the epithelium does not get overheated while the heatis being “pumped” into the deeper lying tissues. Therefore, the serialperiod (t_(ser)) should be longer than about 10 t_(exp), preferablylonger than about 50 t_(exp), and most preferably longer than about 150t_(exp). On the other hand, in order that the deeper lying tissues donot cool down appreciably during the time span between two ESTART laserpulses, the serial period t_(ser) should be shorter than about 3 s,preferably shorter than 1 s, most preferably shorter than 0.5 s.Further, the total duration of the DMR pulse train t_(DMR)=N×t_(ser)should be shorter than 30 s, preferably shorter than 10 s, mostpreferably shorter than 5 s.

An example of the heat being pumped deeper into the tissue during aconsecutive delivery of ESTART pulses is shown in FIG. 11 for the caseof N=20 laser pulses, wherein δ=3 μm, F_(o)=0.8 J/cm² and t_(ser)=50 ms.The white lines represent isothermal curves with marked temperatures in° C.

Considering that the maximal temperature increase during each ESTARTlaser pulse cannot be higher than T_(b), this determines the lower limitfor the required number of pulses as N_(min)=F_(e)/F_(abl).

According to the present invention, one of the preferred embodiments isthe delivery of ESTART* pulses, i.e., pulses satisfying T_(crit)≥T_(b).In this embodiment, ultimate safety can be achieved in a simple mannerby controlling the number N of delivered ESTART* pulses, without theneed for a precise control of the individual laser fluence F_(o) which,as has been shown earlier, can vary uncontrollably during a treatment.Considering that the maximal temperature increase during each ESTART*laser pulse cannot be higher than T_(b), this defines the maximal finaltemperature T_(p1)(see FIG. 10 ) as T_(p1)=T_(o)+N×ΔT_(s) whereΔT_(s)=k(T_(b)−T_(o)). Since ΔT_(s) is determined by the laserparameters t_(ser) and t_(exp) (further determined by δ and t_(ed)),which can be set by means of the control box, the desired maximalsurface temperature T_(p1) can be set by selectingN≤(T_(p1)−T_(o))/(k(T_(b)−T_(o))) at the control box. This maximalsurface temperature T_(p1) holds under any fluence condition. Forexample, for t_(ser)=150 t_(exp) (and therefore k=0.01), we determinethat, by setting N≤13, T_(p1) will never exceed 65° C. (regardless ofF_(o)).

It should be appreciated that, for laser spot dimensions d≤1.5 mm, thethermal diffusion in the lateral direction(s) during the time spant_(ser) can be appreciable, which contributes to the rate of temperaturedecay during the time span between two laser pulses. Therefore, for d imm, the times t_(2%), t_(2%) or t_(3%) will be shorter than previouslydescribed, namely t′_(1%)<75 t_(exp), t′_(2%)<35 t_(exp), and t′_(3%)<8t_(exp), correspondingly. Thus, using a small spot size allows thedelivery of higher fluences at shorter serial times t_(ser). For thisreason, according to one of the preferred embodiments of presentinvention, the energy is delivered to the tissue in a “patterned” shape,wherein the laser beam irradiates a number (M) of individual spots Swithin the treatment area S’ Each spot S having the size (e.g.,diameter) d is separated from a neighbouring spot by the distance x (seeFIG. 12 ). The spot size d and the distance x are chosen such that thespot size d is in the range of 0.3 mm≤d≤1.5 mm, and that the tissuecoverage TC=(M×area(S))/area(S′) (in %) is in the range of 25%≤TC≤65%.Further, the size (e.g., diameter) of the treatment area S′ whichcomprise all the spots is in the range of 3 to 15 mm. These parametersensure that, during the time span between two ESTART pulses (i.e.,during t_(ser)), the thermal diffusion in the lateral direction spreadsthe heat which is generated by the laser radiation away from a localizedspot S towards the surroundings of the spot, thus effectively spatiallyhomogenizing the slowly varying temperature T_(p1) across the area S′.Thus, intense heat shocks with short exposure time can be delivered tothe localized spots S, without causing long-term overheating of thetreatment area S′. This way, it is possible to create an intense heat“needling” on about 50% of the surface of the tissue, whereas thethermal injury in deeper layers is approximately homogeneous (in thelateral direction), since, in the deeper layers, the thermal diffusionspreads the heat in all directions.

There are several possibilities how a plurality of spots within thetreatment area can be generated. First, one can use a screen withvarious holes in front of the laser handpiece. However, if the fluenceof laser radiation becomes too large, the screen which blocks the laserradiation can become too hot. Alternatively, an array of lenses can beused for creating a pattern of distinct spots within the treatment area,wherein the laser beam illuminates the array of lenses. It should benoted that the array of lenses can comprise one layer of lenses whichare transversally arranged with respect to the optical axis or, thearray of lenses can comprise more than one layer of lenses which followeach other on the optical axis. Finally, diffraction optics can be used,in order to generate the pattern of distinct spots within the treatmentarea by using interference effects.

g) Treatment of the Male or Female Urethra by ESTART Pulses

Another aspect of the present invention concerns the treatment of maleor female urinary symptoms and male erectile dysfunction by irradiatingthe urethra with ESTART laser pulses, i.e. laser pulses which have theabove-described advantageous properties. The term erectile dysfunction,as used herein, refers to the inability or impaired ability of a malepatient to experience penile erection. Apart from this, other conditionsarising from degenerative changes of penile connective and vascularnetwork can be treated by the method and the apparatus which isdescribed below.

The term urinary symptoms as used herein refers to at least one of thefollowing symptoms: incontinence (i.e., involuntary leakage), dysuria(i.e., painful urination), urgency and frequency of urination, andrecurrent infections.

g1) Background Information

Erectile dysfunction, i.e., impairment of the ability to achieve ormaintain an erection, is a fairly common medical condition which affectsup to 40% men above the age of 60 and has a detrimental impact onquality of life.

Reasons for erectile dysfunction (ED) are different: they can be ofpsychological, hormonal, neurological or vascular origin. In older men,the main causes of ED are different impairments of vascular function,while, in younger men, other causes are more prevalent. The presentinvention concentrates on the vascular origin of erectile dysfunction.

The body of the penis is composed of three erectile columns, namely twolateral masses which are named “corpora cavernosa” and one median masswhich is named “corpus spongiosum”. The corpus spongiosum contains theurethra in the middle (see FIG. 1 ). The erectile tissues are boundtogether by fascial layers. The erectile tissues consist of connectedcompartments or sinusoids (also called trabeculae) which are divided bythin fibrous membrane, consisting mainly of connective tissue and smoothmuscle fibres.

The penile arterial system is important for supplying the erectiletissues with blood, while the superficial, the intermediate and the deepvenous systems are responsible for the venous drainage of the penis. Theerection of the penis is a neurovascular mechanism, as neurologicstimuli result in a smooth muscle cell relaxation in the cavernoustissue, thus permitting the influx of arterial blood. At the same time,the venous outflow of blood is blocked due to the compression of thevenous network against the fibrous envelope of the penis. This mechanismensures that the erection is maintained.

Erection can be impaired by the degeneration of any of theabove-mentioned structures and tissues which are important for themaintenance of the erection. In the case of vascular malfunction whereeither arterial inflow of blood or venous outflow is affected, problemswith achieving and/or maintaining erection can occur. An importantfactor in the erection mechanism is also the fibrous “skeleton” whichcomprises the erectile tissues and the intra-corporeal fibrous networkas well as smooth muscle cells which enable the relaxation of sinusoidsand the influx of arterial blood.

The therapies for treating erectile dysfunctionality includepsychological, pharmacological or therapies using energy based medicaldevices (the “EBM” devices). EBM devices employ light (laser light orotherwise, as for example the intense pulse light), radiofrequency (RF)radiation, ultrasound or other types of nonchemical energy. In thefollowing, therapies which use laser devices are considered. In order tobe effective, the laser has to deliver enough energy and/or power tothermally affect the structure or the function of the tissues.

Apart from erectile dysfunction, the male or female urinary symptomsrepresent another debilitating condition. The reason is oftenage-related degeneration (atrophy) of the urethral mucosal seal, whichleads to the unwanted leakage of urine and other symptoms.

The present invention involves a laser-based therapy for stimulation andregeneration of the urethral and surrounding tissues in order to treatmale erectile dysfunction, urinary symptoms, and other conditionsarising from degenerative changes of urethral and surrounding (includingpenile) connective and vascular network. In what follows the term“erectile dysfunction” will be used occasionally to represent allconditions arising from degenerative changes of urethral and surrounding(including penile) connective and vascular network. Similarly, the term“penile tissue” will be occasionally used to represent both male andfemale urethral and surrounding tissues.

Most of the previous therapies involve delivering a certain type of anon-chemical energy extra-corporeally to the penile tissue. One of theearlier EBM devices that have been proposed for the stimulation of thehemodynamic activity within the penis is the ultrasound device(WO9912514). The method involves coupling an ultrasound device to theouter surface of the penis and then transmitting the ultrasound energyinto the penis.

Delivering laser radiation intra-urethrally to reverse degenerativechanges of penile connective and vascular network has been avoided,primarily because of the concerns that the locally delivered laserenergy will be transformed into heat which damages the urethra. Duringthe transurethral microwave (i.e., radiofrequency) therapy of theprostate, the energy is delivered through the urethra but not to theurethra. According to this therapy, an instrument (called an antenna)which sends out microwave energy is inserted through the urethra to alocation inside the prostate, wherein the microwave energy is used toheat up the inside of the prostate. In particular, it is the goal thatthat the temperature inside the prostate becomes high enough to killsome of the tissue. During the transurethral microwave therapy specialcautions are taken to prevent that too much heat damages the wall of theurethra, namely by circulating a cooling fluid around the microwaveantenna. In addition, to prevent the temperature from getting too highoutside the prostate, a temperature sensor is inserted into thepatient's rectum during the therapy. If the temperature in the rectumincreases too much, the treatment is turned off until the temperaturefalls again.

g2) Delivering the ESTART Pulses Inside the Urethra

The above-mentioned discoveries have prompted the inventors of thepresent application to use laser thermotherapy to treat conditions thatarise from degenerative changes in the urethral and surrounding tissues,such as the penile urethral and erectile tissues which lead to varioushealth problems, the most serious being erectile dysfunction, maleurinary incontinence, urethral stenosis, and other urinary symptoms.

In an embodiment of the present invention, an intra-urethral cannula isinserted into the urethral opening. Here, a handpiece guiding the laserbeam is connected to the cannula. Alternatively, the cannula is insertedinto the urethral opening first, and then the handpiece is inserted intothe cannula. It should be noted that the handpiece receives the laserbeam from a laser device which generates the laser beam.

The head of cannula is positioned at that region of the urethra whichshall be treated. This way, the laser pulses are delivered to the regionof the urethra which shall be treated. As the laser pulse is absorbed inthe upper layer of the urethra (the epithelium), the laser energy istransferred into heat. The pulse energy is strictly controlled so thattransient pulses of increased temperature are generated in the urethralepithelia. Further, the wavelength and the pulse shape of the laserpulse are selected so that the penetration depth inside the epitheliumis sufficiently small and the energy delivery time of the laser pulse issufficiently short (this means: the laser pulses are ESTART pulses). Inparticular, the penetration depth δ should be less than or equal to 10μm provided that the energy delivery time of the laser pulse is lessthan or equal to 50 μs. Alternatively, for a penetration depth less thanor equal to 4 μm, the energy delivery time should be less than or equalto 250 μs, preferably less than or equal to 150 μs. Finally, for apenetration depth less than or equal to 1 μm the energy exposure timeshould be less than or equal to 350 μs, preferably less than or equal to200 μs.

It should be noted that the cannula together with the handpiece can beslowly pulled out of the urethra so that large sections of the urethraand even the entire urethra can be treated by means of laser pulses.

By using ESTART pulses with an appropriate wavelength and pulse shape,no or only minimal thermal damage is caused in the urethral epithelialsurface, if the fluence (in J/cm²) for the delivered energy to theurethral epithelial surface is appropriately chosen. A particular broadrange of “safe” fluences is possible, if the critical temperatureT_(crit) of the tissue (for the corresponding thermal exposure time ofthe tissue surface) is higher than the tissue boiling temperature T_(b).As described above, the temperature of the tissue cannot rise above thetemperature T_(b) in this case, so that no irreversible damage to thetissue occurs. In this context, it should be noted that the internalsurface of the urethra is not regularly shaped. Also, the handpiece maynot be always kept at the optimal distance or optimal angle with respectto the epithelial surface. This may result in a non-uniform delivery ofthe fluence, and consequently in a non-uniform generation of heat in theradial and longitudinal direction of the urethral cavity. And finally, aphysician error can also occur. In all these cases, the confined boilingmechanism protects the patient from any irreversible damage and hencerepresents a significant safety feature.

Optionally, the section of urethra which needs to be penetrated duringthe laser treatment can be inspected before the laser treatment, i.e.,before the cannula and the handpiece are inserted into the urethra.Here, it should be noted that, as mentioned, the male urethra is curvedand significantly longer than the female urethra so that there is adanger to damage it during a laser treatment. Thus, it is advantageousto observe the urethral region which lies in front of the cannula whenpenetrating the urethra by means of a cannula. This is possible byintroducing an endoscope into the cannula, wherein any endoscope whichhas a smaller external tube diameter than the opening of the cannula canbe used. Thus, it is the cannula with the inserted endoscope that isfirst slowly inserted into the urethra, wherein the endoscope enablesthe operator to constantly monitor the introduction of the cannula intothe urethra. Once the cannula is placed at the appropriate depth, theendoscope is removed, and the handpiece (more precisely, the end of thehandpiece which guides the waveguide) is introduced into the cannula.

According to the present laser treatment of the urethra, thermal energyis deposited throughout the urethra by retrieving the cannula (togetherwith the handpiece) outwards while at the same time emitting the laserpulses. The resulting intense thermal pulsing within the epithelialtissue leads to an intense signalling between the epithelia and theunderlying connective tissues which causes the deeper lying tissues toinitiate a regeneration process. Additionally, the superficial intenseepithelial triggering may be combined with the conventional slowertemperature build-up within the connective tissue (cf. theabove-described dual mechanism regeneration (DMR) treatment). This isaccomplished by delivering a series of short, sharp thermal pulses tothe epithelial tissue, wherein the individual thermal pulses aretemporarily separated by the serial period, t_(ser) which is less thanabout 50 Hz, preferably less than about 16 Hz. In particular, theabove-described dual mechanism regeneration (DMR) treatment which usespulse trains can be also applied to the treatment of the urethra.

This way, a mild hyperthermia is induced in the urethra and theperiurethral erectile tissue (corpus spongiosum) which leads to the ofnew vessels, a regeneration of connective support and an improvement ofthe vascular function. In one of the preferred embodiment of theinvention, the method comprises an additional step which is external andcan be performed simultaneously with intra-urethral laser treatment,namely delivering thermal energy to the anterior surface of the penis.This delivery of thermal energy leads heating pulses for the other twoerectile tissue compartments, the two corpora cavernosa. Thus, theresulting effect of the therapy is the regeneration of the vasculartissue and the fibro-connective support in all tissue compartments whichare together most responsible for the maintenance of an erection.

g3) Apparatus for Treating Urinary Symptoms and Erectile Dysfunction byMeans of Laser Pules

FIG. 13 illustrates applicator elements which can be used for performingthe above-described treatment method of the urethra by means of laserpulses. The applicator elements comprise handpiece 1, cannula 2 andendoscope 3.

As described above, cannula 2 may be first inserted into the urethra,and then handpiece 1 is inserted into cannula 2. Similarly, for theinspection of the urethra, endoscope 3 is inserted into cannula 2. Thus,the handpiece and the endoscope are designed in a manner that they canbe inserted into the cannula, wherein the final position ofhandpiece/endoscope is optionally locked (e.g., screwed) to the cannula.It should be noted that the two applicator elements cannula 2 andhandpiece 1 are an embodiment of the means for introducing a laser pulseinto the urethra according to the present invention. Further, the twoapplicator elements cannula 2 and endoscope 3 are an embodiment of themeans for inspecting the urethra according to the present invention.

Turning to the handpiece 1 in more detail, the thicker part of handpiece1 contains a focusing lens which focuses the laser beam into a hollowwaveguide, which is included in the thinner part of the handpiece. Atthe end of the hollow waveguide, the laser beam then spreads out of thehollow waveguide at an angle of 5-30 degree, preferably 5-15 degree,particularly preferably ca. 11 degree. Since the inner side of thehollow waveguide (which is inserted into the cannula 2 when preformingthe treatment) must not contact any liquid or small particles, cannula 2is covered with a sapphire transparent protective window which protectsthe hollow waveguide of the handpiece 1.

Cannula 2 which is hollow should have an outer diameter so that thecanal of the urethra is widened in a smooth manner. This way, the laserbeam can evenly irradiate the first few millimeters of urethra wallwhich are located in front of the cannula. As mentioned, anotherfunction of the cannula is the protection of the hollow waveguide tipand the endoscope during the treatment. After the treatment, only thecannula needs to be cleaned and sterilized. Besides, cannula 2 comprisesan engraved scale which enables the operator to move the cannula alongthe urethra in a controlled manner.

In an example, a laser device according to the invention may comprise atleast one laser system for generating at least one laser beam, and atleast one optical delivery system for the generated at least laser beam.Particularly, the laser device may comprise two integrated individuallaser systems, each having an individual laser source. The laser devicemay further comprise a control unit, similarly as described withreference to FIG. 5 , for controlling the operation of the at least onelaser source of the at least one laser system, including generated laserbeam parameters. The control unit may control the operation of bothlaser sources and is therefore integral part of both laser systems.However, each laser system may also have its own control unit. It shouldbe noted that the described laser device may be an embodiment of themeans for generating at least one laser pulse comprising a wavelengthaccording to the present invention.

In a preferred embodiment, an optical delivery system includes anarticulated arm and a manually guided laser treatment head or handpiece,as described further above, which is connected to the distal end of thearticulated arm, wherein the laser light is transmitted, relayed,delivered, and/or guided from either one or both laser systems throughthe articulated arm and the laser treatment head to a target.Preferably, the articulated arm can be an Optoflex articulated armavailable from Fotona, d.o.o. (Slovenia, EU). In a preferred embodiment,additionally or alternatively, an optical delivery system may beprovided, wherein, instead of the articulated arm, a flexible elongateddelivery fiber for guiding the laser beam from either one or both lasersystems is incorporated. Either one of the optical delivery systemsmight be used in connection with either one of the two laser systems,thereby guiding either one or both of the laser beams provided by thetwo laser systems. A handpiece as described further above may also beattached to the distal end of the elongated fiber.

Alternatively, one or both laser sources may be built into the handpieceMoreover, the control unit, or the complete laser systems may be builtinto the handpiece as well.

According to the invention, the wavelength of the laser beam which isgenerated by the first laser system is in a range from 2.5 μm to 3.5 μm,preferably 2.7 μm to 3 μm. The wavelength of the laser beam generated bythe second laser system is in the range from 0.8 μm to 11 μm, preferably9 μm to 11 μm. Examples for suitable laser systems are an Er:YAG laser(λ=2.94 mm) or a CO₂ laser (λ=9 to 11 mm).

The invention claimed is:
 1. An apparatus for non-ablative tissueregeneration comprising: a laser for generating laser pulses having awavelength and a pulse energy; a control unit for controlling the laser;and a cannula for delivering the laser pulses onto a surface of aurethra of a human body; and wherein the control unit comprises controlinstructions configured to control the laser so that a delivery timet_(ed) of the laser pulses, during which the second half of the laserpulse energy is delivered, is sufficiently short so that, given thewavelength and thus a corresponding urethra penetration depth ð of thelaser pulses, t_(ed)+(1/D)(δ+√{square root over (2 D t_(ed))})² isshorter than 900 microseconds, wherein D=0.1 mm²s⁻¹; and wherein thecontrol instructions are configured to control the laser such that a sumof the pulse energies of the laser pulses heats the surface of theurethra to a maximal temperature T_(max) between 70° C. and a urethraboiling temperature T_(b).
 2. The apparatus according to claim 1,wherein t_(ed)+(1/D)(δ+√{square root over (2 D t_(ed))})² corresponds toa time interval in which the temperature of the surface of the urethrais above T₀(T_(max)—T₀)/2, wherein To defines an initial temperature ofthe surface of the urethra before the laser pulses are delivered ontothe surface of the urethra, wherein T_(max) defines a maximaltemperature of the surface of the urethra resulting from the deliveredlaser pulses.
 3. The apparatus according to claim 1, wherein the controlinstructions are configured to select the wavelength of the laser pulsesso that the penetration depth δ of the laser pulses are smaller than 30micrometers, preferably smaller than 10 micrometers.
 4. The apparatusaccording to claim 3, wherein the control instructions are configured toselect the wavelength of the laser pulses between 2.6 and 3.2micrometers or between 9.1 and 10.2 micrometers.
 5. The apparatusaccording to claim 1, wherein the control instructions are configured tocontrol the laser so that the energy delivery time t_(ed) of the laserpulses is shorter than 600 microseconds.
 6. The apparatus according toclaim 1, wherein, for t_(ed)+(1/D)(δ+√{square root over (2 D t_(ed))})²being shorter than 900 microseconds, the treated urethra has a criticaltemperature T_(crit) which is greater or equal to the urethra boilingtemperature T_(b); and wherein the control instructions are configuredto select the pulse energy of the laser pulses so large that thecorresponding fluence on the surface of the urethra is up to 4 times theablation threshold fluence F_(abl) of the urethra.
 7. The apparatusaccording to claim 1, wherein, for t_(ed)+(1/D)(δ+√{square root over (2D t_(ed))})² being shorter than 900 microseconds, the treated urethrahas a critical temperature T_(crit) which is smaller than the urethraboiling temperature T_(b); and wherein the control instructions areconfigured to select the pulse energy of the laser pulses sufficientlysmall so that the surface of the urethra is heated to a maximaltemperature T_(max) which is smaller than the critical temperatureT_(crit) of the urethra.
 8. The apparatus according to claim 1, whereinthe control instructions are configured to generate a plurality of laserpulses which are directed to the surface of the urethra; and wherein thecontrol instructions are configured to select the time t_(ser) betweentwo successive laser pulses longer than 10 times t_(ed)+(1/D)(δ+√{squareroot over (2 D t_(ed))})², preferably longer than 40 timest_(ed)+(1/D)(δ+√{square root over (2 D t_(ed))})².
 9. The apparatusaccording to claim 8, wherein the control instructions are configured toselect the time t_(ser) between two successive laser pulses shorter than3 seconds, preferably shorter than 1 second.
 10. The apparatus accordingto claim 8, wherein the control instructions are configured to selectthe number of laser pulses so that the total duration of the pulse trainis shorter than 30 seconds.
 11. The apparatus according to claim 1,wherein the cannula is adapted so that the laser pulses generate two ormore spots on the surface of the urethra, preferably with a spot size din the range 0.3 mm≤d≤1.5 mm.
 12. A method for non-ablative tissueregeneration which uses a laser system and comprises the following step;directing, with a cannula, laser pulses comprising a wavelength and apulse energy onto a surface of a urethra of a human body; wherein theenergy delivery time t_(ed) of the laser pulses, during which the secondhalf of the laser pulse energy is delivered, is chosen sufficientlyshort, so that, given the wavelength and thus a corresponding urethrapenetration depth 5 of the laser pulses, t_(ed)+(1/D)(δ+√{square rootover (2 D t_(ed))})² is smaller than 900 microseconds, wherein D=0.1mm²s⁻¹; and wherein the laser is controlled such that a sum of the pulseenergies of the laser pulses heat the surface of the urethra to amaximal temperature T_(max) between 70° C. and a urethra boilingtemperature T_(b).